The present invention relates generally to magnetic resonance (MR) imaging and, more particularly, to a method and apparatus of acquiring MR data from a motion tracker and gating acquisition of free-breathing MR images.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, or “longitudinal magnetization”, MZ, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt. A signal is emitted by the excited spins after the excitation signal B1 is terminated and this signal may be received and processed to form an image.
When utilizing these signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
Magnetic resonance imaging is frequently used to acquire MR images of the cardiac and abdominal regions of a patient. A concern faced with cardiac and abdominal imaging is the impact of patient breathing. That is, patient breathing often contributes to motion artifacts in the image. As a result, a number of MR scans are conducted during intervals of patient breath-holds. By acquiring data during breath-holds, patient motion can be reduced and, therefore, artifacts reduced.
Requiring patient breath-holding during an MR scan has a number of drawbacks. For example, extensive and repetitive breath-hold moments can lead to patient discomfort and fatigue, in addition to an increasing inability to maintain further breath-holding. Also, breath-holding is generally patient dependent and therefore the physical condition and limitations of a given patient may limit the MR scan. In other words, breath-holding time limits the spatial resolution and number of slices that can be acquired. This can be particularly problematic for 3D imaging. That is, patients are often unable to suspend breathing for the time necessary for routine 3D scans.
Thus, as an alternative to breath-holding, respiratory gating techniques have been developed to prospectively or retrospectively gate the acquisition of MR data such that free-breathing can occur during data acquisition. One such gating technique is generally referred to navigator echo imaging. This technique is used to directly detect the position of the right hemi-diaphragm and uses the positional information to gate data acquisition from a free-breathing patient. Specifically, a spin echo sequence with perpendicularly intersecting slab planes or a cylindrical excitation gradient echo sequence is used to selectively excite a column of spins. The position of the column is generally perpendicular to the dome of the right hemi-diaphragm, the anterior left ventricle wall, or other locations where respiratory or other motion can be captured. Post-processing of the resulting echo yields the position of the tracked object, which can be used as a basis for navigator gating.
A conventional navigator echo pulse sequence is illustrated in FIG. 1. Specifically, FIG. 1 illustrates the sequence diagram of a conventional electrocardiogram (ECG) gated, navigator gated, fat suppressed (saturation), segmented 3D FGRE/FIESTA sequence. As illustrated, the pulse sequence is triggered by an ECG trigger signal 10 that is defined by R peaks 12. The time interval between the R-wave peaks define the periodic cardiac (R-R) interval. With this conventional MR pulse sequence, data is acquired after a period of cardiac delay. Following the period of cardiac delay, a navigator excitation pulse 14 is played out to sample a navigator echo. This navigator excitation is followed by a fat saturation segment 16. Thereafter, a segmented part of k-space is acquired during imaging acquisition segment 18. These segments 14, 16, and 18 repeat each heart cycle, i.e., each R-R interval, until all the k-space views have been acquired.
In the case of whole heart imaging or liver imaging, an axial imaging slab is generally defined across the entire heart or liver and the navigator tracker is defined at the right hemi-diaphragm of the patient. This conventional placement, however, poses a problem for quality of the navigator tracker profile because it is partially overlapped by the axial imaging slab, as illustrated in FIG. 2. During the first heart cycle, after the navigator saturation pulse and the fat saturation pulses have been applied 20 and 22, respectively, the k-space segment acquisition for the first heart cycle is acquired 24. After this acquisition segment is complete, some of the spins within the navigator tracker column will be saturated by the excitation RF pulse defining the imaging slab whereas other spins will not. Line 26 corresponds to the magnetization of the overlapped spins whereas line 28 corresponds to the non-overlapped spins. As all the spins recover at the same rated based on their longitudinal magnetization left at the end of the k-space segment (M′z), at the beginning of the acquisition at the next heart cycle 30, the spins within the navigator tracker column will be recovered to:Mz(tn)=M′zexp(−tn/T1)+Mo(1−exp(−tn/T1))  (Eqn. 1),
where Mz(tn) is the longitudinal magnetization at time tn, tn is the time from the navigator saturation RF excitation pulse to the navigator saturation RF excitation pulse of the next heart cycle, M′z is the longitudinal magnetization present immediately after the image acquisition, M0 is the equilibrium longitudinal magnetization, and T1 is the spin-lattice relaxation time of the tissue/blood. As the residual longitudinal magnetization (M′z) is lower in the area overlapped by the imaging slab, the signal intensity in those areas (proportional to Mz(t)) will be reduced and the shape of the navigator profile will be distorted. Most motion detection techniques such as edge detection, least-squares error minimization, and cross-correlation are sensitive to the shape of the profile. Therefore, the motion detection accuracy will be reduced and the image quality of the resulting image will be degraded by motion artifact. FIG. 3 illustrates the navigator spatial profile 32 at the second heart cycle. As illustrated, the reduction in longitudinal magnetization in spins within the imaging slab results in a significant signal intensity drop at the area intersected by the imaging slab 34.
It would therefore be desirable to have a system and method capable of navigator echo tracking with patient free-breathing that is less susceptible to the saturation effect on the shape of the signal intensity profile described above.